System and method for determining a temperature during analyte measurement

ABSTRACT

A method of measuring an analyte in a biological fluid comprises applying an excitation signal having a DC component and an AC component. The AC and DC responses are measured; a corrected DC response is determined using the AC response; and a concentration of the analyte is determined based upon the corrected DC response. Other methods and devices are disclosed.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a divisional of U.S. patent application Ser. No.10/687,668 filed Oct. 17, 2003, which is a continuation-in-part of U.S.patent application Ser. No. 10/264,890, filed Oct. 4, 2002, which is adivisional application of U.S. patent application Ser. No. 09/530,171,filed Apr. 24, 2000, which is the U.S. national stage of InternationalPatent Application Serial No. PCT/US98/27203, filed Dec. 21, 1998, whichis a continuation-in-part of U.S. patent application Ser. No.08/996,280, filed Dec. 22, 1997 (now abandoned), and claims prioritybenefit from each of these applications. This application also claimsthe benefit of U.S. Provisional Application No. 60/480,298, filed Jun.20, 2003. The contents of each of these applications are herebyincorporated by reference herein.

TECHNICAL FIELD OF THE INVENTION

The present invention relates to a measurement method and apparatus foruse in measuring concentrations of an analyte in a fluid. The inventionrelates more particularly, but not exclusively, to a method andapparatus which may be used for measuring the concentration of glucosein blood.

BACKGROUND OF THE INVENTION

Measuring the concentration of substances, particularly in the presenceof other, confounding substances, is important in many fields, andespecially in medical diagnosis. For example, the measurement of glucosein body fluids, such as blood, is crucial to the effective treatment ofdiabetes.

Diabetic therapy typically involves two types of insulin treatment:basal, and meal-time. Basal insulin refers to continuous, e.g.time-released insulin, often taken before bed. Meal-time insulintreatment provides additional doses of faster acting insulin to regulatefluctuations in blood glucose caused by a variety of factors, includingthe metabolization of sugars and carbohydrates. Proper regulation ofblood glucose fluctuations requires accurate measurement of theconcentration of glucose in the blood. Failure to do so can produceextreme complications, including blindness and loss of circulation inthe extremities, which can ultimately deprive the diabetic of use of hisor her fingers, hands, feet, etc.

Multiple methods are known for measuring the concentration of analytesin a blood sample, such as, for example, glucose. Such methods typicallyfall into one of two categories: optical methods and electrochemicalmethods. Optical methods generally involve reflectance or absorbancespectroscopy to observe the spectrum shift in a reagent. Such shifts arecaused by a chemical reaction that produces a color change indicative ofthe concentration of the analyte. Electrochemical methods generallyinvolve, alternatively, amperometric or coulometric responses indicativeof the concentration of the analyte. See, for example, U.S. Pat. No.4,233,029 to Columbus, U.S. Pat. No. 4,225,410 to Pace, U.S. Pat. No.4,323,536 to Columbus, U.S. Pat. No. 4,008,448 to Muggli, U.S. Pat. No.4,654,197 to Lilja et al., U.S. Pat. No. 5,108,564 to Szuminsky et al.,U.S. Pat. No. 5,120,420 to Nankai et al., U.S. Pat. No. 5,128,015 toSzuminsky et al., U.S. Pat. No. 5,243,516 to White, U.S. Pat. No.5,437,999 to Diebold et al., U.S. Pat. No. 5,288,636 to Pollmann et al.,U.S. Pat. No. 5,628,890 to Carter et al., U.S. Pat. No. 5,682,884 toHill et al., U.S. Pat. No. 5,727,548 to Hill et al., U.S. Pat. No.5,997,817 to Crismore et al., U.S. Pat. No. 6,004,441 to Fujiwara etal., U.S. Pat. No. 4,919,770 to Priedel, et al., and U.S. Pat. No.6,054,039 to Shieh, which are hereby incorporated in their entireties.

An important limitation of electrochemical methods of measuring theconcentration of a chemical in blood is the effect of confoundingvariables on the diffusion of analyte and the various active ingredientsof the reagent. For example, the geometry and state of the blood samplemust correspond closely to that upon which the signal-to-concentrationmapping function is based.

The geometry of the blood sample is typically controlled by asample-receiving portion of the testing apparatus. In the case of bloodglucose meters, for example, the blood sample is typically placed onto adisposable test strip that plugs into the meter. The test strip may havea sample chamber (capillary fill space) to define the geometry of thesample. Alternatively, the effects of sample geometry may be limited byassuring an effectively infinite sample size. For example, theelectrodes used for measuring the analyte may be spaced closely enoughso that a drop of blood on the test strip extends substantially beyondthe electrodes in all directions. Ensuring adequate coverage of themeasurement electrodes by the sample, however, is an important factor inachieving accurate test results. This has proven to be problematic inthe past, particularly with the use of capillary fill spaces.

Other examples of limitations to the accuracy of blood glucosemeasurements include variations in blood composition or state (otherthan the aspect being measured). For example, variations in hematocrit(concentration of red blood cells), or in the concentration of otherchemicals in the blood, can effect the signal generation of a bloodsample. Variations in the temperature of blood samples is yet anotherexample of a confounding variable in measuring blood chemistry.

Thus, a system and method are needed that accurately measure bloodglucose, even in the presence of confounding variables, includingvariations in temperature, hematocrit, and the concentrations of otherchemicals in the blood. A system and method are also needed to ensureadequate coverage of the measurement electrodes by the sample,particularly in capillary fill devices. A system and method are likewiseneeded that accurately measure an analyte in a fluid. It is an object ofthe present invention to provide such a system and method.

SUMMARY OF THE INVENTION

In one embodiment of the present invention, a method for measuring atemperature during a test for determining a concentration of a medicallysignificant component of a biological fluid is disclosed, comprising thesteps of a) providing a sensor having a reaction zone in which thebiological fluid reacts with a reagent; and measuring a temperature inthe reaction zone.

In another embodiment of the present invention, a method for measuring atemperature during a test for determining a concentration of a medicallysignificant component of a biological fluid is disclosed, comprising thesteps of a) applying a first signal having an AC component to thebiological fluid; b) measuring a first AC response to the first signal;and using the first AC response to produce an indication of atemperature of the biological fluid.

In yet another embodiment of the present invention, a method ofaccounting for the effect of a temperature variation on a test for aglucose concentration of a biological fluid is disclosed comprising: a)applying at least a first test signal having an AC component to thebiological fluid, the first test signal having a first frequency; b)measuring at least a first AC response to the first test signal; c)determining a temperature value of the biological fluid using the firstAC response; and d) determining a temperature-corrected glucoseconcentration of the biological fluid based at least in part upon thetemperature value.

In another embodiment of the present invention, a method of accountingfor the effect of a temperature variation on a test for a glucoseconcentration of a biological fluid is disclosed comprising a) applyingat least a first test signal having an AC component to the biologicalfluid, the first test signal having a first frequency; b) measuring atleast a first AC response to the first test signal; c) determining atemperature value of the biological fluid using the first AC response;and d) determining a temperature-corrected glucose concentration of thebiological fluid based at least in part upon the temperature value.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention will be further described, by way of example only, withreference to the accompanying drawings, in which:

FIG. 1 is a diagram of a first embodiment excitation signal suitable foruse in a system and method according to the present invention, having aserially-applied AC component and DC component.

FIG. 2 is a diagram of a second embodiment excitation signal suitablefor use in a system and method according to the present invention,having a simultaneously-applied AC component and DC component.

FIGS. 3A-B illustrate a first embodiment test strip of the presentinvention.

FIG. 4 is a diagram of an excitation signal utilized in the test ofExample 1.

FIG. 5 is a plot of the correlation coefficient r² (glucose vs. DCcurrent) versus Read Time for the test of Example 1 with no incubationtime.

FIG. 6 is a plot of the correlation coefficient r² (glucose vs. DCcurrent) versus Read Time for the test of Example 1 with varyingincubation time.

FIG. 7 is a plot of AC admittance versus hematocrit for the test ofExample 2.

FIG. 8 is a plot of uncompensated DC current versus glucose for the testof Example 2.

FIG. 9 is a plot of the predicted glucose response versus the actualglucose response for the test of Example 2.

FIG. 10 is a diagram of an excitation signal utilized in the test ofExample 3.

FIG. 11 is a plot of the AC phase angle versus reference glucose for thetest of Example 3.

FIG. 12 is a plot of the predicted glucose response versus the actualglucose response for the test of Example 3.

FIG. 13 is a diagram of an excitation signal utilized in the test ofExample 4.

FIG. 14 is a plot of AC admittance versus hematocrit (parametricallydisplayed with temperature) for the test of Example 4.

FIG. 15 is a plot of the uncompensated DC response versus actual glucosefor the test of Example 4.

FIG. 16 is a plot of the predicted glucose response versus actualglucose response for the test of Example 4.

FIGS. 17A-B illustrate a second embodiment test strip of the presentinvention.

FIG. 18 is a plot parametrically illustrating the correlationcoefficient r² between the DC current response and glucose level as ReadTime varies for three combinations of temperature and hematocrit in thetest of Example 5.

FIG. 19 is a diagram of the excitation signal utilized in the test ofExample 5.

FIG. 20 is a plot of AC admittance versus hematocrit as temperature isparametrically varied in the test of Example 5.

FIG. 21 is a plot of AC admittance phase angle versus hematocrit astemperature is parametrically varied in the test of Example 5.

FIG. 22 is a plot of the uncompensated DC response versus actual glucosefor the test of Example 5.

FIG. 23 is a plot of the predicted glucose response versus actualglucose response for the test of Example 5.

FIG. 24 is a diagram of the excitation signal utilized in the test ofExample 6.

FIG. 25 is a plot of the correlation coefficient r² between hematocritand DC response current plotted against hematocrit in the test ofExample 6.

FIG. 26 is a plot of AC admittance phase angle versus hematocrit for thetest of Example 6.

FIG. 27 is a plot of the uncompensated DC response versus actual glucosefor the test of Example 6.

FIG. 28 is a plot of the compensated DC response versus actual glucosefor a 1.1 second Total Test Time of Example 6.

FIG. 29 is a plot of the compensated DC response versus actual glucosefor a 1.5 second Total Test Time of Example 6.

FIG. 30 is a plot of the compensated DC response versus actual glucosefor a 1.9 second Total Test Time of Example 6.

FIG. 31 is a table detailing the heights and widths of the capillaryfill channels used in the test devices of Example 8, as well asschematic diagrams of convex and concave sample flow fronts in acapillary fill space.

FIGS. 32A-C are schematic plan views of a test strip illustrating thepotential for biased measurement results when a concave flow frontencounters a prior art dose sufficiency electrode.

FIG. 33 is a schematic plan view of a test strip of the presentinvention having a pair of perpendicular dose sufficiency electrodesthat are independent from the measurement electrodes.

FIGS. 34A-B are schematic plan views of the test strip of FIG. 33containing samples with convex and concave flow fronts, respectively.

FIGS. 35A-B are schematic plan views of a test strip of the presentinvention having a pair of parallel dose sufficiency electrodes that areindependent from the measurement electrodes.

FIG. 36 is a schematic plan view of the test strip of FIG. 35,schematically illustrating the electric field lines that communicatebetween the electrode gap when the electrodes are covered with sample.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

For the purposes of promoting an understanding of the principles of theinvention, reference will now be made to the embodiment illustrated inthe drawings, and specific language will be used to describe thatembodiment. It will nevertheless be understood that no limitation of thescope of the invention is intended. Alterations and modifications in theillustrated device, and further applications of the principles of theinvention as illustrated therein, as would normally occur to one skilledin the art to which the invention relates are contemplated, are desiredto be protected. In particular, although the invention is discussed interms of a blood glucose meter, it is contemplated that the inventioncan be used with devices for measuring other analytes and other sampletypes. Such alternative embodiments require certain adaptations to theembodiments discussed herein that would be obvious to those skilled inthe art.

The entire disclosure of U.S. provisional applications titled DEVICESAND METHODS RELATING TO ELECTROCHEMICAL BIOSENSORS (Ser. No. 60/480,243,filed Jun. 20, 2003) and DEVICES AND METHODS RELATING TO ANALYTE SENSOR(Ser. No. 60/480,397, Filed Jun. 20, 2003) are hereby incorporated byreference in their entireties.

A system and method according to the present invention permit theaccurate measurement of an analyte in a fluid. In particular, themeasurement of the analyte remains accurate despite the presence ofinterferants, which would otherwise cause error. For example, a bloodglucose meter according to the present invention measures theconcentration of blood glucose without error that is typically caused byvariations in the temperature and the hematocrit level of the sample.The accurate measurement of blood glucose is invaluable to theprevention of blindness, loss of circulation, and other complications ofinadequate regulation of blood glucose in diabetics. An additionaladvantage of a system and method according to the present invention isthat measurements can be made much more rapidly and with much smallersample volumes, making it more convenient for the diabetic person tomeasure their blood glucose. Likewise, accurate and rapid measurement ofother analytes in blood, urine, or other biological fluids provides forimproved diagnosis and treatment of a wide range of medical conditions.

It will be appreciated that electrochemical blood glucose meterstypically (but not always) measure the electrochemical response of ablood sample in the presence of a reagent. The reagent reacts with theglucose to produce charge carriers that are not otherwise present inblood. Consequently, the electrochemical response of the blood in thepresence of a given signal is intended to be primarily dependent uponthe concentration of blood glucose. Secondarily, however, theelectrochemical response of the blood to a given signal is dependentupon other factors, including hematocrit and temperature. See, forexample, U.S. Pat. Nos. 5,243,516; 5,288,636; 5,352,351; 5,385,846; and5,508,171, which discuss the confounding effects of hematocrit on themeasurement of blood glucose, and which are hereby incorporated byreference in their entireties. In addition, certain other chemicals caninfluence the transfer of charge carriers through a blood sample,including, for example, uric acid, bilirubin, and oxygen, therebycausing error in the measurement of glucose.

A preferred embodiment system and method for measuring blood glucoseaccording to the present invention operates generally by using thesignal-dependence of the contribution of various factors to theimpedance (from which admittance and phase angle may be derived) of ablood sample. Because the contribution of various factors to theimpedance of a blood sample is a function of the applied signal, theeffects of confounding factors (that is, those other than the factorssought to be measured) can be substantially reduced by measuring theimpedance of the blood sample to multiple signals. In particular, theeffects of confounding factors, (primarily temperature and hematocrit,but also including chemical interferants such as oxygen), contributeprimarily to the resistivity of the sample, while the glucose-dependentreaction contributes primarily to the capacitance. Thus, the effects ofthe confounding factors can be eliminated by measuring the impedance ofthe blood sample to an AC excitation, either alone or in combinationwith a DC excitation. The impedance (or the impedance derived admittanceand phase information) of the AC signal is then used to correct the DCsignal or AC derived capacitance for the effects of interferants.

It will be appreciated that measurements at sufficiently high ACfrequencies are relatively insensitive to the capacitive component ofthe sample's impedance, while low frequency (including DC) measurementsare increasingly (with decreasing frequency) sensitive to both theresistive and the capacitive components of the sample's impedance. Theresistive and capacitive components of the impedance can be betterisolated by measuring the impedance at a larger number of frequencies.However, the cost and complexity of the meter increases as the number ofmeasurements increases and the number of frequencies that need to begenerated increases. Thus, in the presently preferred embodiment, theimpedance may be measured at greater than ten frequencies, butpreferably at between two and ten frequencies, and most preferably atbetween two and five frequencies.

As used herein, the phrase “a signal having an AC component” refers to asignal which has some alternating potential (voltage) portions. Forexample, the signal may be an “AC signal” having 100% alternatingpotential (voltage) and no DC portions; the signal may have AC and DCportions separated in time; or the signal may be AC with a DC offset (ACand DC signals superimposed).

Sample Measurement with Successive AC and DC Signals

FIG. 1 illustrates a preferred embodiment excitation signal suitable foruse in a system and method according to the present invention, indicatedgenerally at 100, in which DC excitation and four frequencies of ACexcitation are used. FIG. 1 also illustrates a typical response to theexcitation when the excitation is applied to a sample of whole bloodmixed with an appropriate reagent, the response indicated generally at102. A relatively high frequency signal is applied, starting at time101. In the preferred embodiment the frequency is between about 10 kHzand about 20 kHz, and has an amplitude between about 12.4 mV and about56.6 mV. A frequency of 20 kHz is used in the example of FIG. 1. Thoseskilled in the art will appreciate that these values may be optimised tovarious parameters such as cell geometry and the particular cellchemistry.

At time 110 a test strip is inserted into the meter and several possibleresponses to the insertion of the test strip into the glucose meter areshown. It will be appreciated that the test strip may also be insertedbefore the excitation signal 100 is initiated (i.e. before time 101);however, the test strip itself may advantageously be tested as a controlfor the suitability of the strip. It is therefore desirable that theexcitation signal 100 be initiated prior to test strip insertion. Forexample, relatively large current leakage, as shown at 112, may occur ifthe strip is wet, either because the test strip was pre-dosed, or due toenvironmental moisture. If the test strip has been pre-dosed andpermitted to largely or completely dry out, an intermediate currentleakage may occur, as shown at 114. Ideally, insertion of the test stripwill cause no or negligible leakage current due to an expected absenceof charge carriers between the test electrodes, as shown at 116.Measured current leakage above a predetermined threshold level willpreferably cause an error message to be displayed and prevent the testfrom continuing.

Once a suitable test strip has been inserted, the user doses the strip,as shown at time 120. While the blood sample is covering the electrodesthe current response will rapidly increase, as the glucose reacts withthe reagent and the contact area increases to maximum. The responsecurrent will reach a stable state, which indicates the impedance of thesample at this frequency. Once this measurement is made and recorded bythe test meter, the excitation frequency is then stepped down to about10 kHz in the preferred embodiment, as shown at time 130. Anothermeasurement is made and recorded by the test meter, and the frequency isstepped down to about 2 kHz in the preferred embodiment, as shown at140. A third measurement is made and recorded by the test meter at thisfrequency. A fourth measurement is made at about 1 kHz in the preferredembodiment, as shown at 150. In the preferred embodiment, measurementsare taken at regular intervals (e.g. 10 points per cycle). It will beappreciated that the stable state response may be measured as current orvoltage (preferably both magnitude and phase) and the impedance and/oradmittance can be calculated therefrom. Although the presentspecification and claims may refer alternately to the AC response asimpedance or admittance (magnitude and/or phase), resistance,conductivity, current or charge, and to the DC response as current,charge, resistance or conductivity, those skilled in the art willrecognize that these measures are interchangeable, it only beingnecessary to adjust the measurement and correction mathematics toaccount for which measure is being employed. In the preferredembodiment, the test meter applies a voltage to one electrode andmeasures the current response at the other electrode to obtain both theAC and DC response.

In certain alternative embodiments measurements are made at fewer ormore frequencies. Preferably measurements are made at at least two ACfrequencies at least an order of magnitude apart. If more than two ACfrequencies are used, then it is preferable that the highest and lowestfrequencies be at least an order of magnitude apart.

It will be appreciated that various waveforms may be used in an ACsignal, including, for example, sinusoidal, trapezoidal, triangle,square and filtered square. In the presently preferred embodiment the ACsignal has a filtered square waveform that approximates a sine wave.This waveform can be generated more economically than a true sine wave,using a square wave generator and one or more filters.

Once all four AC measurements are made, the signal is preferably brieflyreduced to zero amplitude, as shown at 160. The DC excitation is thenbegun, as shown at 170. The amplitude of the DC excitation isadvantageously selected based on the reagent being used, in order tomaximise the resulting response or response robustness. For example, ifferricyanide is being used in a biamperometry system, the DC amplitudeis preferably about 300 mV. For another example, if a nitrosoanilinederivative is being used in a biamperometry system, the DC amplitude ispreferably about 500-550 mV. In the alternative, if a third referenceelectrode is used, the DC applitude is preferably 600 mV (versus thesilver/silver chloride reference electrode) for ferricyanide, and 40-100mV (versus the silver/silver chloride reference electrode) fornitrosoaniline derivative. During DC excitation, measurements arepreferably made at a rate of 100 pts/sec. The current response willfollow a decay curve (known as a Cottrell curve), as the reaction islimited by the diffusion of unreacted glucose next to the workingelectrode. The resulting stable-state amplitude (measured or projected)is used to determine a glucose estimation of the sample, as is known inthe art. A corrected estimation is then determined that corresponds moreclosely to the concentration of glucose in the blood, by using theimpedance of the sample to the AC signal to correct for the effects ofinterferants, as explained in greater detail hereinbelow.

It will be appreciated that a method according to the present inventionmay also be used to measure the concentration of other analytes and inother fluids. For example, a method according to the present inventionmay be used to measure the concentration of a medically significantanalyte in urine, saliva, spinal fluid, etc. Likewise, by appropriateselection of reagent a method according to the present invention may beadapted to measure the concentration of, for example, lactic acid,hydroxybutyric acid, etc.

Sample Measurement with Simultaneously Applied AC and DC Signals

It will be appreciated that at least some of the applied DC and ACcomponents can also be applied simultaneously. FIG. 2 illustrates anexcitation signal suitable for use in a system and method according tothe present invention in which some of the AC and DC components areapplied simultaneously, indicated generally at 200, and havingcorresponding events numbered correspondingly to FIG. 1 (so, forexample, the signal 200 is initiated at time 201, and a strip isinserted at time 210, etc.). As with the signal 100, the signal 200 hasa frequency of about 10-20 kHz and an amplitude of about 12.4-56.6 mV.However, after the strip has been dosed, as shown at time 220, a DCoffset is superimposed, as shown at 270. Typical AC and DC responses areshown in FIG. 2. The AC and DC responses are measured simultaneously andmathematically deconvoluted and used to determine the impedance(admittance magnitude and phase) and the amperometric or coulometricresponse.

A system for measuring blood glucose according to the present inventionadvantageously employs a blood glucose meter and test strips generallysimilar to those used in prior art systems, such as those commerciallyavailable from Roche Diagnostics, and such as are described in U.S. Pat.Nos. 6,270,637; and 5,989,917, which are hereby incorporated in theirentireties. These test strips provide apparati having a sample cell inwhich the blood sample is received for testing, and electrodes disposedwithin the sample cell through which the excitation signal is providedand the measurements are made. Those skilled in the art will appreciatethat these test strips and meters may advantageously be used for themeasurement of glucose in blood, but that other apparati may be moresuitable for the measurement of other analytes or other biologicalfluids when practising the present invention.

A suitable glucose meter may be adapted from such known meters by theaddition of electronic circuitry that generates and measures signalshaving AC and DC components, such as those described hereinabove, and bybeing programmed to correct the DC measurement using the ACmeasurement(s), as described in greater detail hereinbelow. It will beappreciated that the specific geometry and chemistry of the test stripscan cause variations in the relationships between the concentration ofglucose, hematocrit, and temperature, and the impedance of a sample.Thus, a given combination of test strip geometry and chemistry must becalibrated, and the meter programmed with the corresponding algorithm.The present invention comprehends the application of excitation signalsin any order and combination. For example, the present inventioncomprehends the application of 1) AC only, 2) AC then DC, 3) AC then DCthen AC, 4) DC then AC, and 5) AC with a DC offset, just to name a fewof the possible permutations.

The use of the complex AC impedance measurement data to correct for theeffects of interferants on the DC measurement is advantageouslyillustrated by the following series of examples. These examplesillustrate how the principles of the present invention can facilitateimprovements in accuracy and test speed when measuring the concentrationof an analyte in a test specimen. Although the following examples dealwith correcting for the interfering effects of hematocrit andtemperature on blood glucose determinations, those skilled in the artwill recognize that the teachings of the present invention are equallyuseful for correcting for the effects of other interferants in bothblood glucose measurements and in the measurement of other analytes.Furthermore, the present specification and claims refer to steps such as“determine the hematocrit value” and “determine the temperature,” etc.To use the hematocrit value as an example, it is intended that suchstatements include not only determining the actual hematocrit value, butalso a hematocrit correction factor vs. some nominal point. In otherwords, the process may never actually arrive at a number equal to thehematocrit value of the sample, but instead determine that the sample'shematocrit differs from a nominal value by a certain amount. Bothconcepts are intended to be covered by statements such as “determine thehematocrit value.”

EXAMPLE 1 DC-Only Measurement Dose Response Study

The measurements made in Example 1 were achieved using the test stripillustrated in FIGS. 3A-B and indicated generally at 300. The test strip300 includes a capillary fill space containing a relatively thick filmreagent and working and counter electrodes, as described in U.S. Pat.No. 5,997,817, which is hereby incorporated by reference. The test strip300 is commercially available from Roche Diagnostics Corporation(Indianapolis, Ind.) under the brand name Comfort Curve®. Theferricyanide reagent used had the composition described in Tables I andII. TABLE I Reagent Mass Composition - Prior to Dispense and Drying Massfor Component % w/w 1 kg solid Polyethylene oxide (300 kDa) 0.8400%8.4000 g solid Natrosol 250M 0.0450% 0.4500 g solid Avicel RC-591F0.5600% 5.6000 g solid Monobasic potassium phosphate 1.2078% 12.0776 g(annhydrous) solid Dibasic potassium phosphate 2.1333% 21.3327 g(annhydrous) solid Sodium Succinate hexahydrate 0.6210% 6.2097 g solidQuinoprotein glucose 0.1756% 1.7562 g dehydrogenase (EnzC#: 1.1.99.17)solid PQQ 0.0013% 0.0125 g solid Trehalose 2.0000% 20.0000 g solidPotassium Ferricyanide 5.9080% 59.0800 g solid Triton X-100 0.0350%0.3500 g solvent Water 86.4731% 864.7313 g % Solids 0.1352687 Target pH6.8 Specific Enzyme Activity Used (U/mg) 689 DCIP Dispense Volume perSensor 4.6 mg

TABLE II Reagent Layer Composition - After Drying Mass per Component %w/w Sensor solid Polyethylene oxide (300 kDa) 6.2099% 38.6400 ug solidNatrosol 250M 0.3327% 2.0700 ug solid Avicel RC-591F 4.1399% 25.7600 ugsolid Monobasic potassium phosphate 8.9286% 55.5568 ug (annhydrous)solid Dibasic potassium phosphate 15.7706% 98.1304 ug (annhydrous) solidSodium Succinate hexahydrate 4.5906% 28.5646 ug solid Quinoproteinglucose dehydrogenase 1.2983% 8.0784 ug (EnzC#: 1.1.99.17) solid PQQ0.0093% 0.0576 ug solid Trehalose 14.7854% 92.0000 ug solid PotassiumFerricyanide 43.6760% 271.7680 ug solid Triton X-100 0.2587% 1.6100 ug

In the measurements, blood samples were applied to test strip 300 andthe excitation potentials illustrated in FIG. 4 were applied to theelectrodes. The excitation comprised a 2 kHz 40 mV_(rms) (56.56 mV peak)AC signal applied between 0 seconds and approximately 4.5 seconds aftersample application, followed by a 300 mV DC signal applied thereafter.For the calculations of this example, however, only the DC measurementdata was analyzed.

In order to determine the minimum needed DC excitation time, a “doseresponse” study was performed, in which glycollyzed (glucose depleted)blood was divided into discrete aliquots and controlled levels ofglucose were added to obtain five different known levels of glucose inthe blood samples. The resulting DC current profile was then examined astwo parameters were varied. The first parameter was the Incubation Time,or the time between the detection of the blood sample being applied tothe test strip 300 and the application of the DC potential to the teststrip 300. The second parameter to be varied was the Read Time, or thetime period after application of the DC potential and the measurement ofthe resulting current. The length of time between detection of the bloodsample being applied to the test strip to the taking of the lastmeasurement used in the concentration determination calculations is theTotal Test Time. In this study, therefore, the sum of the IncubationTime and the Read Time is the Total Test Time. The results of this studyare illustrated in FIGS. 5 and 6.

In FIG. 5, the DC response was measured with no incubation time (ReadTime=Total Test Time). FIG. 5 plots the correlation coefficient r²versus Read Time. As can be seen, the correlation exceeds 0.95 within1.0 second. In FIG. 6, the DC response was measured with varyingIncubation Time. When an Incubation Time is provided (even an IncubationTime as short as two (2) seconds), the r² value rose to over 0.99 in 0.5seconds or less after application of the DC potential.

The barrier to implementation of such fast test times in a consumerglucose test device, however, is the variation from blood sample toblood sample of the level of interference from the presence of bloodcells in the sample. The hematocrit (the percentage of the volume of ablood sample which is comprised of cells versus plasma) varies fromindividual to individual. The interference effect of hematocrit on suchmeasurements is fairly complex. In the tests of Example 1, however, allsamples contained the same level of hematocrit. With no variablehematocrit influence at the different glucose levels, the hematocritterm cancels out in the correlation figures.

EXAMPLE 2 Combined AC and DC Measurement of Capillary Blood Samples

The measurements made in Example 2 were also achieved using the teststrip illustrated in FIGS. 3A-B and indicated generally at 300. Asdescribed above, the test strip 300 includes a capillary fill spacecontaining a relatively thick film reagent and working and counterelectrodes, as described in U.S. Pat. No. 5,997,817, which is herebyincorporated herein by reference.

In the measurements, capillary blood samples from various fingerstickdonors were applied to test strip 300 and the excitation potentialsillustrated in FIG. 4 were applied to the electrodes. The excitationcomprised a 2 kHz 40 mV_(rms) AC signal applied between 0 seconds andapproximately 4.5 seconds after sample application, followed by a 300 mVDC signal applied thereafter.

In this Example 2, the AC response of the sample was derived asadmittance (the inverse of impedance). The admittance response isproportionate to the hematocrit level of the sample in a temperaturedependent manner. The relationship between admittance, hematocrit andtesting temperature is illustrated in FIG. 7. The data used for theadmittance charted in FIG. 7 is the last admittance measurement made foreach sample during the AC portion of the excitation illustrated in FIG.4.

Regression analysis of this data allows admittance, hematocrit andtemperature to be related according to the following formula:H _(est) =c ₀ +c ₁ Y _(2 kHz) +c ₂ dT   (Equation 1)Using this relationship to predict the blood hematocrit is accomplishedusing test temperature data reported by the temperature sensor in themeter and the measured admittance. In Equation 1, c₀, c₁ and c₂ areconstants, dT is the deviation in temperature from a center defined as“nominal” (24° C. for example), and H_(est) is the estimated deviationin hematocrit from a similar “nominal” value. For the present purposes,the actual hematocrit value is not necessary, and it is generallypreferred to produce a response which is proportionate but centersaround a nominal hematocrit. Thus, for a 70% hematocrit, the deviationfrom a nominal value of 42% would be 28%, while conversely for a 20%hematocrit the deviation from that same nominal value would be −22%.

By using the AC admittance measurement to estimate the hematocrit levelusing Equation 1, the accuracy of the DC glucose response can be greatlyimproved by combining the estimated hematocrit, temperature and DCresponse to correct for the hematocrit interference in the DC responseas follows:PRED=(a ₀ +hct ₁ H _(est) +hct ₂ H _(est) ² +tau ₁ dT+tau ₂ dT ²)+(a ₁DC)(1+hct ₃ H _(est) +hct ₄ H _(est)2)(1+tau ₃ dT+tau ₄ dT ²)  (Equation 2)where DC is the measured glucose current response to the applied DCsignal and PRED is the compensated (predicted) glucose responsecorrected for the effects of hematocrit and temperature. The constants(a₀, hct₁, hct₂ , tau₁, tau₂, a₁, hct₃, hct₄, tau₃ and tau₄) in Equation2 can be determined using regression analysis, as is known in the art.

FIG. 8 illustrates the uncompensated 5.5 second DC glucose response ofall of the capillary blood samples as temperature varies (ignoring theAC measurement data). As will be appreciated, there is a wide variationin the DC current response as temperature and hematocrit vary. FIG. 9illustrates the correlation between the actual blood glucose level ofthe sample versus the predicted response using Equation 2. As can beseen, when the DC response is compensated for hematocrit levels usingthe AC response data, r² values of 0.9404 to 0.9605 are achieved with aTotal Test Time of 5.5 seconds.

EXAMPLE 3 Use of AC Phase Angle to Estimate Blood Glucose Levels andHematocrit

The measurements made in Example 3 were also achieved using the teststrip illustrated in FIGS. 3A-B and indicated generally at 300. Asdescribed above, the test strip 300 includes a capillary fill spacecontaining a relatively thick film reagent and working and counterelectrodes, as described in U.S. Pat. No. 5,997,817, which is herebyincorporated by reference. Because hematocrit levels from capillaryblood samples typically vary only between 30%-50%, spiked venous bloodsamples having a hematocrit range from 20%-70% were used for thisExample 3. Five levels of glucose, temperature (14, 21, 27, 36 and 42°C.) and hematocrit (20, 30, 45, 60 and 70%) were independently varied,producing a covariance study with 125 samples.

In the measurements, blood samples were applied to test strip 300 andthe excitation potentials illustrated in FIG. 10 were applied to theelectrodes. The excitation comprised a 2 kHz AC signal for approximately4.1 seconds, a 1 kHz AC signal for approximately 0.1 seconds, and a 200Hz signal for approximately 0.1 seconds. All three AC signals had anamplitude of 56.56 mV peak. No DC excitation was used in this example.The Total Test Time was 4.3 seconds from sample application time.

It was found that another component of the AC response, the phase angle(particularly at lower frequencies, such as 200 Hz in this Example 3),is also a function of the sample glucose level in the case of this teststrip and reagent. This relationship is demonstrated in FIG. 11, wherethe AC phase angle for each of the three test frequencies is plottedversus the reference glucose level. Regression analysis for each of thethree frequencies produces AC phase angle-to-reference glucose level r²correlation values of 0.9114 at 2 kHz, 0.9354 at 1 kHz, and 0.9635 at200 Hz. The present invention therefore comprehends the use of the ACphase angle to measure glucose levels. The AC excitation frequencyproducing the measured phase angle is preferably 2 kHz or below, morepreferably 1 kHz or below, and most preferably 200 Hz or below, but notincluding DC excitation.

The linearized relationship between the 200 Hz phase angle response andthe blood glucose level is as follows:P _(eff)=(Φ_(200 Hz)/Γ)^(−γ)  (Equation 3)where P_(eff) is the effective phase, which is proportional to glucose,the terms Γ and γ are constants, and Φis the measured AC phase angle.

Using the same approach to compensate for temperature and hematocrit asused in Example I above (see Equations 1 and 2) produced a predictivealgorithm as follows:PRED=(a ₀ +hct ₁ H _(est) +hct ₂ H _(est) ² +tau ₁ dT+tau ₂ dT ²)+(a ₁ P_(off))(1+hct ₃ H _(est) +hct ₄ H _(est)2)(1+tau ₃ dT+tau ₄ dT ²)  (Equation 4)The resulting compensated (predicted) response PRED versus glucose forthe 125 blood samples (each tested with eight test strips) is shown inFIG. 12. The r² correlation of the PRED response vs. known glucoselevel, where all temperatures and all hematocrits are combined, is0.9870. This Example 3 demonstrates again the value of AC measurementsfor compensating for interferants that reduce the accuracy of bloodglucose measurements. Using an existing commercially available sensor,the present invention yields a 4.3 second Total Test Time with anoverall r² of 0.9870.

It was also determined that AC phase angle measurements can producehematocrit level measurements that are almost immune to the effects oftemperature variation. In another covariant study of 125 samples (fiveglucose concentrations, five hematocrit concentrations and fivetemperatures), each of the samples was tested using an excitationprofile of 20 kHz, 10 kHz, 2 kHz, 1 kHz and DC. The AC phase angle atvarious frequencies was related to glucose, hematocrit and temperatureusing linear regression to determine the coefficients of the followingformula at each of the four AC frequencies:Phase=c₀+c₁ Glu+c ₂ HCT+c ₃Temp   (Equation 5)where Glu is the known glucose concentration, HCT is the knownhematocrit concentration and Temp is the known temperature.

The determined coefficients revealed that the temperature coefficient(C₃) was essentially zero at 20 kHz and 10 kHz, cancelling temperaturefrom the equation at these frequencies. Furthermore, the glucosecoefficient (c₁) is essentially zero at all of the AC frequenciesbecause, as explained hereinabove, the higher frequency AC impedancemeasurements are largely unaffected by glucose levels and are thereforeuseful for measuring the levels of interfering substances. It wastherefore found that the hematocrit level could be determinedindependent of temperature and glucose level using only the AC phaseangle measurements. In a preferred embodiment, the hematocrit may bemeasured using the phase angle data from all four measured frequencies:H _(est) =c ₀ +c ₁Φ_(20 kHz) +c ₂Φ_(10 kHz) +c ₃Φ_(2 kHz) +hd 4Φ_(1 kHz)   (Equation 6)Those skilled in the art will recognise that that the coefficients canbe empirically determined for any particular test strip architecture andreagent chemistry. The present invention therefore may be used toestimate hematocrit using only AC phase angle measurements preferablymade at at least one AC frequency, more preferably made at at least twoAC frequencies, and most preferably made at at least four ACfrequencies.

EXAMPLE 4 Combined AC and DC Measurement Using Nitrosoaniline Reagent

The measurements made in Example 4 were also achieved using the teststrip illustrated in FIGS. 3A-B and indicated generally at 300. Asdescribed above, the test strip 300 includes a capillary fill spacecontaining a relatively thick film reagent and working and counterelectrodes, as described in U.S. Pat. No. 5,997,817, which is herebyincorporated by reference. The test strip was modified from thatdescribed in U.S. Pat. No. 5,997,817, however, by the use of a differentreagent. The nitrosoaniline reagent used had the composition describedin Tables III and IV. TABLE III Reagent Mass Composition - Prior toDispense and Drying Mass for Component % w/w 1 kg solid Polyethyleneoxide (300 kDa) 0.8054% 8.0539 g solid Natrosol 250M 0.0470% 0.4698 gsolid Avicel RC-591F 0.5410% 5.4104 g solid Monobasic potassiumphosphate 1.1437% 11.4371 g (annhydrous) solid Dibasic potassiumphosphate 1.5437% 15.4367 g (annhydrous) solid Disodium Succinatehexahydrate 0.5876% 5.8761 g solid Potassium Hydroxide 0.3358% 3.3579 gsolid Quinoprotein glucose 0.1646% 1.6464 g dehydrogenase (EnzC#:1.1.99.17) solid PQQ 0.0042% 0.0423 g solid Trehalose 1.8875% 18.8746 gsolid Mediator 31.1144 0.6636% 6.6363 g solid Triton X-100 0.0327%0.3274 g solvent Water 92.2389% 922.3888 g % Solids 0.1352687 Target pH6.8 Specific Enzyme Activity Used (U/mg) 689 DCIP Dispense Volume perSensor 4.6 mg

TABLE IV Reagent Layer Composition - After Drying Mass per Component %w/w Sensor solid Polyethylene oxide (300 kDa) 10.3829% 37.0480 ug solidNatrosol 250M 0.6057% 2.1611 ug solid Avicel RC-591F 6.9749% 24.8877 ugsolid Monobasic potassium phosphate 14.7445% 52.6107 ug (annhydrous)solid Dibasic potassium phosphate 19.9006% 71.0087 ug (annhydrous) solidDisodium Succinate hexahydrate 7.5753% 27.0299 ug solid PotassiumHydroxide 4.3289% 15.4462 ug solid Quinoprotein glucose dehydrogenase2.1225% 7.5734 ug (EnzC#: 1.1.99.17) solid PQQ 0.0546% 0.1947 ug solidTrehalose 24.3328% 86.8243 ug solid Mediator BM 31.1144 8.5553% 30.5268ug solid Triton X-100 0.4220% 1.5059 ug

The method for the manufacture of the glucose biosensor for this Example4 is the same in all respects as disclosed in U.S. Pat. No. 5,997,817except for the manufacture of the reagent. A protocol for thepreparation of the preferred embodiment nitrosoaniline reagent is asfollows:

Step 1: Prepare a buffer solution by adding 1.54 g of dibasic potassiumphosphate (anhydrous) to 43.5 g of deionized water. Mix until thepotassium phosphate is dissolved.

Step 2: To the solution from step 1, add 1.14 g of monobasic potassiumphosphate and mix until dissolved.

Step 3: To the solution from step 2, add 0.59 g of disodium succinate(hexahydrate) and mix until dissolved.

Step 4: Verify that the pH of the solution from step 3 is 6.7±0.1.Adjustment should not be necessary.

Step 5: Prepare a 5 g aliquot of the solution from step 4, and to thisadd 113 kilounits (by DCIP assay) of the apoenzyme of quinoproteinglucose dehydrogenase (EC#: 1.1.99.17). This is approximately 0.1646 g.Mix, slowly, until the protein is dissolved.

Step 6: To the solution from step 5, add 4.2 milligrams of PQQ and mixfor no less than 2 hours to allow the PQQ and the apoenzyme toreassociate in order to provide functional enzyme.

Step 7: To the solution from step 4, add 0.66 g of the mediatorprecursor, N,N-bis(hydroxyethyl)-3-methoxy-4-nitrosoaniline(hydrochloride) (BM 31.1144). Mix until dissolved (this solution willhave a greenish black coloration).

Step 8: Measure the pH of the solution from step 7 and adjust the pH toa target of 7.0±0.1. Normally this is accomplished with 1.197 g of SNpotassium hydroxide. Because the specific amount of potassium hydroxidemay vary as needed to reach the desired pH, generally deviations in massfrom the 1.1 97 g are made up from an aliquot of 3.309 g deionized waterwhich is also added at this step.

Step 9: Prepare a solution of Natrosol 250M (available from Aqualon), byslowly sprinkling 0.047 g over 44.57 g of deionized water which is mixed(using a rotary mixer and blade impeller) at a rate of approximately 600rpm in a vessel of sufficient depth such that the rotor blades are notexposed nor the solution running over. Mix until the Natrosol iscompletely dissolved.

Step 10: Prepare a suspension of Avicel RC-591F (available from FMS), byslowly sprinkling 0.54 g onto the surface of the solution from step 9,mixing at a rate of approximately 600 rpm for not less than 60 minutesbefore proceeding.

Step 11: To the suspension from step 10, gradually add 0.81 g ofPolyethylene oxide of 300 kDa mean molecular weight while mixing andcontinue to mix for not less than 60 minutes before proceeding.

Step 12: Gradually add the solution from step 8 to the suspension fromstep 11 while mixing. Reduce the mixing rate to 400 rpm.

Step 13: To the reagent from step 12, add 1.89 g of Trehalose andcontinue mixing for not less than 15 minutes.

Step 14: To the reagent from step 13, add 32.7mg of Triton X-100(available from Roche Diagnostics) and continue mixing.

Step 15: To the reagent from step 14, add the enzyme solution from step6. Mix for no less than 30 minutes. At this point the reagent iscomplete. At room teperature the wet reagent mass is consideredacceptable for use for 24 hours.

Spiked venous blood samples were used. Five levels of glucose, fourtemperatures (19, 23, 32 and 38° C.) and five levels of hematocrit (20,30, 45, 60 and 70%) were independently varied, producing a covariancestudy with 100 samples. 16 test strips 300 were tested for each uniquecombination of glucose, temperature and hematocrit. The blood sampleswere applied to test strip 300 and the excitation potentials illustratedin FIG. 13 were applied to the electrodes. The excitation comprised a3.2 kHz AC signal for approximately 4.0 seconds, a 2.13 kHz AC signalfor approximately 0.1 seconds, a 1.07 kHz AC signal for approximately0.1 seconds, a 200 Hz AC signal for approximately 0.1 seconds, a 25 HzAC signal for approximately 0.1 seconds, followed by a DC signal of 550mV for approximately 1.0 second. All four AC signals had an amplitude of56.56 mV peak. The Total Test Time was 5.5 seconds from sampleapplication time.

In this Example 4, the AC response of the sample was derived asadmittance (the inverse of impedance). The admittance response isproportionate to the hematocrit level of the sample in a temperaturedependent manner. The relationship between admittance, hematocrit andtesting temperature is illustrated in FIG. 14. As compared to the teststrip architecture of Example 2, the orthogonality of the temperatureand hematocrit influence on glucose was not as strong in this Example 4,therefore a cross product term (T×HCT) was added to the admittanceregression formula used in FIG. 14. The data used for the admittancecharted in FIG. 14 is the last admittance measurement made for eachsample during the 3.2 kHz AC portion of the excitation illustrated inFIG. 13.

Regression analysis of this data allows admittance, hematocrit andtemperature to be related according to the following formula:H _(est)=(Y _(3.2 kHz) +c ₀ +c ₁ dT)/(c ₂ dT+c ₃)   (Equation 7)It was determined that the admittance measurement made at 3.2 kHz wasbest correlated with hematocrit for this test system. Using thisrelationship to predict the blood hematocrit is accomplished using testtemperature data reported by the temperature sensor in the meter and themeasured admittance. In Equation 7, c₀, c₁, c₂ and c₃ are constants, dTis the deviation in temperature from a center defined as “nominal” (24°C. for example), and H_(est) is the estimated deviation in hematocritfrom a similar “nominal” value. For the present purposes, the actualhematocrit value is not necessary, and it is generally preferred toproduce a response which is proportionate but centers around a nominalhematocrit. Thus, for a 70% hematocrit, the deviation from a nominalvalue of 42% would be 28%, while conversely for a 20% hematocrit thedeviation from the same nominal value would be −22%.

By using the AC admittance measurement to estimate the hematocrit levelusing Equation 7, the accuracy of the DC glucose response can be greatlyimproved by combining the estimated hematocrit, temperature and DCresponse to correct for the hematocrit interference in the DC responseas follows (same as Equation 2 above):PRED=(a ₀ +hct ₁ H _(est) +hct ₂ H _(est) +tau ₁ dT+tau ₂ dT ²)+(a₁DC)(1+hct ₃H_(est) +hct ₄ H _(est) ²)(1+tau ₃ dT+tau ₄ dT ²)   (Equation8)The constants in Equation 8 can be determined using regression analysis,as is known in the art.

FIG. 15 illustrates the uncompensated 5.5 second DC glucose response ofall of the blood samples as hematocrit and temperature vary (ignoringthe AC measurement data). As will be appreciated, there is a widevariation in the DC current response as temperature and hematocrit vary.FIG. 16 illustrates the correlation between the actual blood glucoselevel of the sample versus the predicted response using Equation 8. Ascan be seen, when the DC response is compensated for hematocrit levelsusing the AC response data, an overall r² value of 0.9818 is achievedwith a Total Test Time of 5.5 seconds. This demonstrates theapplicability of the present invention in achieving high accuracy andfast test times with a different reagent class than was used in Examples1-3.

EXAMPLE 5 Combined AC and DC Measurement Using a 0.397 μl Sample

The measurement methods of the present invention have been found to beuseful with other test strip designs as well. Example 5 was conductedusing the test strip design illustrated in FIGS. 17A-B, and indicatedgenerally at 1700. Referring to FIG. 17A, the test strip 1700 comprisesa bottom foil layer 1702 formed from an opaque piece of 350 μm thickpolyester (in the preferred embodiment this is Melinex 329 availablefrom DuPont) coated with a 50 nm conductive (gold) layer (by sputteringor vapor deposition, for example). Electrodes and connecting traces arethen patterned in the conductive layer by a laser ablation process toform working, counter, and dose sufficiency electrodes (described ingreater detail hereinbelow) as shown. The laser ablation process isperformed by means of an excimer laser which passes through achrome-on-quartz mask. The mask pattern causes parts of the laser fieldto be reflected while allowing other parts of the field to pass through,creating a pattern on the gold which is ejected from the surface wherecontacted by the laser light.

Examples of the use of laser ablation techniques in preparing electrodesfor biosensors are described in U.S. patent application Ser. No.09/866,030, “Biosensors with Laser Ablation Electrodes with a ContinuousCoverlay Channel” filed May 25, 2001, and in U.S. patent applicationSer. No. 09/411,940, entitled “Laser Defined Features for PatternedLaminates and Electrode,” filed Oct. 4, 1999, both disclosuresincorporated herein by reference.

The bottom foil layer 1702 is then coated in the area extending over theelectrodes with a reagent layer 1704 in the form of an extremely thinreagent film. This procedure places a stripe of approximately 7.2millimeters width across the bottom foil 1702 in the region labelled“Reagent Layer” on FIG. 17. In the present Example, this region iscoated at a wet-coat weight of 50 grams per square meter of coatedsurface area leaving a dried reagent less than 20 μm thick. The reagentstripe is dried conventionally with an in-line drying system where thenominal air temperature is at 110° C. The rate of processing isnominally 30-38 meters per minute and depends upon the rheology of thereagent.

The materials are processed in continuous reels such that the electrodepattern is orthogonal to the length of the reel, in the case of thebottom foil 1702. Once the bottom foil 1702 has been coated withreagent, the spacer is slit and placed in a reel-to-reel process ontothe bottom foil 1702. Two spacers 1706 formed from 100 μm polyester (inthe preferred embodiment this is Melinex 329 available from DuPont)coated with 25 μm PSA (hydrophobic adhesive) on both the dorsal andventral surfaces are applied to the bottom foil layer 1702, such thatthe spacers 1706 are separated by 1.5 mm and the working, counter anddose sufficiency electrodes are centered in this gap. A top foil layer1708 formed from 100 μm polyester coated with a hydrophilic film on itsventral surface (using the process described in U.S. Pat. No. 5,997,817) is placed over the spacers 1706. In the preferred embodiment,the hydrophilic film is coated with a mixture of Vitel and Rhodapexsurfactant at a nominal thickness of 10 microns. The top foil layer 1708is laminated using a reel-to-reel process. The sensors can then beproduced from the resulting reels of material by means of slitting andcutting.

The 1.5 mm gap in the spacers 1706 therefore forms a capillary fillspace between the bottom foil layer 1702 and the top foil layer 1708.The hydrophobic adhesive on the spacers 1706 prevents the test samplefrom flowing into the reagent under the spacers 1706, thereby definingthe test chamber volume. Because the test strip 1700 is 5 mm wide andthe combined height of the spacer 1706 and conductive layer is 0.15 mm,the sample receiving chamber volume is5 mm×1.5 mm×0.15 mm=1.125 μl   (Equation 9)

As shown in FIG. 17B, the distance from the sample application port 1710and the dose sufficiency electrodes is 1.765 mm. The volume of sampleneeded to sufficiently cover the working, counter and dose sufficiencyelectrodes (i.e. the minimum sample volume necessary for a measurement)is1.5 mm×1.765 mm×0.15 mm=0.397 μl   (Equation 10)

The reagent composition for the test strip 1700 is given in Tables V andVI. TABLE V Reagent Mass Composition - Prior to Dispense and Drying Massfor Component % w/w 1 kg solid Polyethylene oxide (300 kDa) 1.0086%10.0855 g solid Natrosol 250M 0.3495% 3.4954 g solidCarboxymethylcellulose 7HF 0.3495% 3.4954 g solid Monobasic potassiumphosphate 0.9410% 9.4103 g (annhydrous) solid Dibasic potassiumphosphate 1.6539% 16.5394 g (trihydrous) solid Disodium Succinatehexahydrate 0.2852% 2.8516 g solid Potassium Hydroxide 0.2335% 2.3351 gsolid Quinoprotein glucose 0.3321% 3.3211 g dehydrogenase (EnzC#:1.1.99.17) solid PQQ 0.0093% 0.0925 g solid Trehalose 0.7721% 7.7210 gsolid Mediator 31.1144 0.6896% 6.8956 g solid Triton X-100 0.0342%0.3419 g solvent Water 93.7329% 937.3293 g % Solids 6.6585% Target pH 7Specific Enzyme Activity Used (U/mg) 689 DCIP Wet Reagent Coat Weightper Sensor (ug/mm²) 50

TABLE VI Reagent Layer Composition - After Drying Mass per Component %w/w Sensor* solid Polyethylene oxide (300 kDa) 15.1469% 3.7821 ug solidNatrosol 250M 5.2495% 1.3108 ug solid Carboxymethylcellulose 7HF 5.2495%1.3108 ug solid Monobasic potassium phosphate 14.1328% 3.5289 ug(annhydrous) solid Dibasic potassium phosphate 24.8395% 6.2023 ug(trihydrous) solid Disodium Succinate hexahydrate 4.2827% 1.0694 ugsolid Potassium Hydroxide 3.5069% 0.8757 ug solid Quinoprotein glucosedehydrogenase 4.9878% 1.2454 ug (EnzC#: 1.1.99.17) solid PQQ 0.1390%0.0347 ug solid Trehalose 11.5958% 2.8954 ug solid Mediator BM31.114410.3562% 2.5859 ug solid Triton X-100 0.5135% 0.1282 ug*“Mass per Sensor” is the amount of the component within the capillary;this does not reflect the reagent that is outside of the capillary.

A protocol for the preparation of the preferred embodimentnitrosoaniline reagent is as follows:

Step 1: Prepare a buffer solution by adding 1.654 g of dibasic potassiumphosphate (trihydrous) to 31.394 g of deionized water. Mix until thepotassium phosphate is dissolved.

Step 2: To the solution from step 1, add 0.941 g of monobasic potassiumphosphate and mix until dissolved.

Step 3: To the solution from step 2, add 0.285 g of disodium succinate(hexahydrate) and mix until dissolved.

Step 4: Verify that the pH of the solution from step 3 is 6.8±0.1.Adjustment should not be necessary. Step 5: Prepare a 4.68 g aliquot ofthe solution from step 4, and to this add 229 kilounits (by DCIP assay)of the apoenzyme of quinoprotein glucose dehydrogenase (EC#: 1.1.99.17).This is approximately 0.3321 g. Mix, slowly, until the protein isdissolved.

Step 6: To the solution from step 5, add 9.3 milligrams of PQQ and mixfor no less than 2 hours to allow the PQQ and the apoenzyme toreassociate in order to provide functional enzyme.

Step 7: Prepare a solution by dissolving 0.772 g of Trehalose into 1.218g of deionized water.

Step 8: After enzyme reassociation, add the solution from step 7 to thesolution from step 6 and continue mixing for not less than 30 minutes.

Step 9: To the solution from step 4, add 0.690 g of the mediatorprecursor BM 31.1144. Mix until dissolved (this solution will have agreenish black coloration).

Step 10: Measure the pH of the solution from step 9 and adjust the pH toa target of 7.0±0.1. Normally this is accomplished with 1.006 g of SNpotassium hydroxide. Because the specific amount of potassium hydroxidemay vary as needed to reach the desired pH, generally deviations in massfrom the 1.006 g are made up from an aliquot of 3.767 g deionized waterwhich is also added at this step.

Step 11: Prepare a solution of Natrosol 250M (available from Aqualon),by slowly sprinkling 0.350 g over 56.191 g of deionized water which ismixed (using a rotary mixer and blade impeller) at an initial rate ofapproximately 600 rpm in a vessel of sufficient depth such that therotor blades are not exposed nor the solution running over. As theNatrosol dissolves, the mixing rate needs to be increased to a speed of1.2-1.4 krpm. Mix until the Natrosol is completely dissolved. Note thatthe resulting matrix will be extremely viscous—this is expected.

Step 12: To the solution from step 11, gradually add 0.350 g ofSodium-Carboxymethylcellulose 7HF (available from Aqualon). Mix untilthe polymer is dissolved.

Step 13: To the suspension from step 13, gradually add 1.01 g ofPolyethylene oxide of 300 kDa mean molecular weight while mixing andcontinue to mix for not less than 60 minutes before proceeding.

Step 14: Gradually add the solution from step 10 to the suspension fromstep 13 while mixing.

Step 15: To the reagent from step 14, add 34.2 mg of Triton X-100(available from Roche Diagnostics) and continue mixing.

Step 16: To the reagent from step 15, add the enzyme solution from step8. Mix for no less than 30 minutes. At this point the reagent iscomplete. At room teperature the wet reagent mass is consideredacceptable for use for 24 hours.

The measurement results illustrated in FIG. 18 show the correlationcoefficient r² between the DC current response and the glucose level asthe Read Time varies for three combinations of temperature andhematocrit. These results demonstrate that a robust DC response shouldbe anticipated for tests as fast as 1 second. However, those skilled inthe art will recognise that there are undesirable variations in thesensor accuracy (correlation) due to the interfering effects oftemperature and hematocrit levels, suggesting that the combined AC andDC measurement method of the present invention should produce moreclosely correlated results.

Based upon the encouraging results obtained in FIG. 18, a further testwas designed using the excitation signal of FIG. 19 applied to the teststrip 1700. The excitation comprised a 10 kHz AC signal applied forapproximately 1.8 seconds, a 20 kHz AC signal applied for approximately0.2 seconds, a 2 Hz AC signal applied for approximately 0.2 seconds, a 1Hz AC signal applied for approximately 0.2 seconds, and a DC signalapplied for approximately 0.5 seconds. The AC signals had an amplitudeof 12.7 mV peak, while the DC signal had an amplitude of 550 mV. TheTotal Test Time was 3.0 seconds.

A covariance study using spiked venous blood samples representing fiveglucose levels (40, 120, 200, 400 and 600), five hematocrit levels (20,30, 45, 60 and 70%) and five temperatures (12, 18, 24, 32 and 44° C.)was designed, resulting in 125 separate combinations. As in the previousexamples, the relationship between admittance, temperature andhematocrit was examined and plotted (FIG. 20 shows the admittance at 20kHz versus hematocrit as temperature varies) and it was confirmed thatthe admittance was linearly related to hematocrit in a temperaturedependent manner. An additional discovery, however, was that the phaseangle of the AC response was correlated with hematocrit in a temperatureindependent manner. The phase angle of the 20 kHz AC response is plottedversus hematocrit in FIG. 21. The results for phase angle measured at 10kHz are similar. The hematocrit of the blood sample may therefore bereliably estimated using only the phase angle information as follows:H _(est) =c ₀ +c ₁(Φ_(10 kHz)−Φ_(20 kHz))+c₂(Φ_(2 kHz)−Φ_(1 kHz))  (Equation 11)

For the test strip used in this Example 5, the correlation between phaseangle and hematocrit was better at higher frequencies. Because of this,the c₂ constant approaches zero and H_(est) can reliably be estimatedusing only the 10 kHz and 20 kHz data. Use of lower frequencies,however, allows for slight improvements in the strip-to-stripvariability of the H_(est) function. The present invention therefore maybe used to estimate hematocrit using only AC phase angle measurementspreferably made at at least one AC frequency, more preferably made at atleast two AC frequencies, and most preferably made at at least four ACfrequencies.

Because the hematocrit can be determined using only the AC responsedata, and we know from FIG. 20 that admittance is linearly related tohematocrit and temperature, we can now determine the temperature of thesample under analysis using only the AC response as follows:T _(est) =b ₀ +b ₁(Y _(10 kHz) −Y _(20 kHz))+b ₂(Y _(2 kHz) −Y_(1 kHz))+b ₃ H _(est)   (Equation 12)where b₀, b₁, b₂ and b₃ are constants. It will be appreciated that theestimation of hematocrit and temperature from the AC response data maybe made with more or fewer frequency measurements, and at differentfrequencies than those chosen for this example. The particularfrequencies that produce the most robust results will be determined bytest strip geometries and dimensions. The present invention thereforemay be used to estimate test sample temperature using only AC responsemeasurements preferably made at at least one AC frequency, morepreferably made at at least two AC frequencies, and most preferably madeat at least four AC frequencies.

Those skilled in the art will recognise that the direct measurement ofthe temperature of the sample under test (by means of the AC response)is a great improvement over prior art methods for estimating thetemperature of the sample. Typically, a thermistor is placed in the testmeter near where the test strip is inserted into the meter. Because thethermistor is measuring a temperature remote from the actual sample, itis at best only a rough approximation of the true sample temperature.Furthermore, if the sample temperature is changing (for example due toevaporation), then the thermal inertia of the test meter and even thethermistor itself will prevent the meter-mounted thermistor fromaccurately reflecting the true temperature of the sample under test. Bycontrast, the temperature estimation of the present invention is derivedfrom measurements made within the sample under test (i.e. within thereaction zone in which the sample under test reacts with the reagent),thereby eliminating any error introduced by the sample being remote fromthe measuring location. Additionally, the temperature estimation of thepresent invention is made using data that was collected very close intime to the glucose measurement data that will be corrected using thetemperature estimation, thereby further improving accuracy. Thisrepresents a significant improvement over the prior art methods.

As a demonstration of the effectiveness of the method of this Example 5for correcting for the effects of interferants on the blood glucosemeasurement, the uncompensated DC current response versus known glucoseconcentration is plotted in FIG. 22 for all 125 combinations of glucose,temperature and hematocrit (the AC measurements were ignored whenplotting this data). As will be appreciated by those skilled in the art,the data exhibits huge variation with respect to hematocrit andtemperature.

As previously discussed, the accuracy of the DC glucose response can begreatly improved by combining the estimated hematocrit, temperature andDC response to correct for the hematocrit and temperature interferencein the DC response as follows:PRED=(a ₀ +hct ₁ H _(est) +hct ₂ H _(est) ² +tau ₁ T _(est) +tau ₂ T_(est))+(a ₁ DC)(1+hct ₃ H _(est) +hct ₄ H _(est) ²)(1+tau ₃ T _(est)+tau ₄ T _(est))   (Equation 13)The constants in Equation 13 can be determined using regressionanalysis, as is known in the art. The present invention therefore allowsone to estimate hematocrit by using the AC phase angle response(Equation 11). The estimated hematocrit and the measured AC admittancecan be used to determine the estimated temperature (Equation 12).Finally, the estimated hematocrit and estimated temperature can be usedwith the measured DC response to obtain the predicted glucoseconcentration (Equation 13).

Applying the above methodology to the test data plotted in FIG. 22, weobtain the predicted glucose versus DC current response illustrated inFIG. 23. This data represents 125 covariant samples having hematocritlevels ranging from 20%-70% and temperatures ranging from 12° C.-44° C.Even with these wide variations in interferant levels, the measurementmethod of the present invention produced an overall r² correlation of0.9874 using a 3.0 second Total Test Time.

EXAMPLE 6 Simultaneous AC and DC Measurement Using a 0.397 μl Sample

Using the same test strip 1700 and reagent described above for Example5, the excitation profile illustrated in FIG. 24 was utilized in orderto decrease the Total Test Time. As described above with respect toExample 5, it was determined that the phase angle at 20 kHz and at 10kHz were most closely correlated with the hematocrit estimation. It wastherefore decided to limit the AC portion of the excitation to these twofrequencies in Example 6 in order to decrease the Total Test Time. Inorder to make further reductions in Total Test Time, the 10 kHz ACexcitation was applied simultaneously with the DC signal (i.e. an ACsignal with a DC offset), the theory being that this combined mode wouldallow for the collection of simultaneous results for DC current, ACphase and AC admittance, providing the fastest possible results.Therefore, the 20 kHz signal was applied for 0.9 seconds. Thereafter,the 10 kHz and DC signals were applied simultaneously for 1.0 secondafter a 0.1 second interval.

For this Example 6, 49 spiked venous blood samples representing sevenglucose levels and seven hematocrit levels were tested. The correlationcoefficient r² between the DC current and the blood hematocrit was thenexamined at three DC measurement times: 1.1 seconds, 1.5 seconds and 1.9seconds after sample application. These correlations are plotted versushematocrit level in FIG. 25. All of these results are comparable,although the correlation is generally poorest at 1.1 seconds andgenerally best at 1.5 seconds. The minimum correlation coefficient,however, exceeds 0.99.

FIG. 26 illustrates the phase angle at 20 kHz plotted against hematocritlevels. The correlation between these two sets of data is very good,therefore it was decided that the 10 kHz data was unnecessary forestimating hematocrit. The hematocrit can therefore be estimated solelyfrom the 20 kHz phase angle data as follows:H _(est) =c ₀ +c ₁Φ_(20 kHz)   (Equation 14)

FIG. 27 illustrates the DC current response versus glucose level for allmeasured hematocrit levels as the read time is varied between 1.1seconds, 1.5 seconds and 1.9 seconds. Not surprisingly, the DC currentat 1.1 seconds is greater than the DC current at 1.5 seconds, which isgreater than the DC current at 1.9 seconds. Those skilled in the artwill recognise that the hematocrit level has a large effect on the DCcurrent, particularly at high glucose concentrations.

As discussed hereinabove, the accuracy of the DC glucose response can begreatly improved by compensating for the interference caused byhematocrit as follows:PRED=(a ₀ +hct ₁ H _(est) +hct ₂ H _(est) ²)+(a ₁ DC)(1+hct ₃ H _(est)+hct ₄ H _(est) ²)   (Equation 15)Note that Equation 15 does not include temperature compensation termssince temperature variation was not included in the experiment of thisExample 6, it can be reasonably inferred from previous examples that aTest term could be included using the 10 kHz and 20 kHz admittancevalues in combination with the H_(est) term. Because the hematocrit canbe reliably estimated using only the 20 kHz phase angle measurementdata, the hematocrit compensated predicted glucose response can bedetermined using only this phase angle information and the measured DCresponse. The compensated DC response versus glucose level for only theDC read at 1.1 seconds (representing a 1.1 second Total Test Time) isillustrated in FIG. 28. The data shows an overall r² correlation of0.9947 with a 1.1 second Total Test Time.

The same data for the 1.5 second DC read is illustrated in FIG. 29,showing an overall r² correlation of 0.9932 for a 1.5 second Total TestTime. The same data for the 1.9 second DC read is illustrated in FIG.30, showing an overall r² correlation of 0.9922 for a 1.9 second TotalTest Time. Surprisingly, the r² correlation actually decreased slightlywith the longer test times. Notwithstanding this, the correlationcoefficients for all three compensated data sets—where all 7 hematocritsranging from 20% through 60% are combined—were in excess of 0.99,demonstrating the applicability of the present invention to yield ablood glucose test as fast as 1.1 seconds, combined with improvedaccuracy, where the sensor requires less than 0.4 microliters of bloodin order to perform the glucose measurement test.

EXAMPLE 7 Use of AC Phase Angle to Detect an Abused Sensor

In order to provide an extra measure of quality control to the analytemeasurement process, particularly when the test system is to be used bya non-professional end user, it is desirable to detect sensors (teststrips) that have been mis-dosed (double dosed, etc.), that have beenpreviously used, or that have degraded enzymes (from being stored in toohumid an environment, being too old, etc.). These conditions arecollectively referred to as “abused sensors.” It is desired to devise atest that will abort the analyte measurement process (or at least warnthe user that the test results may not be accurate) if an abused sensoris inserted into the test meter.

When performing a blood glucose analysis, the test meter will typicallymake several successive current measurements as the blood samplecontinues to react with the reagent chemistry. As is well known in theart, this response current is known as the Cottrell current and itfollows a pattern of decay as the reaction progresses. We may define aCottrell Failsafe Ratio (CFR) as follows:

The Cottrell response of the biosensor in the Confidence system can begiven by: $\begin{matrix}{I_{cottrell} = {\frac{{nFA}\sqrt{D}}{\sqrt{\Pi}}{Ct}^{\alpha}}} & \left( {{Equation}\quad 16} \right)\end{matrix}$where: n=electrons freed per glucose molecule

-   -   F=Faraday's Constant    -   A=Working electrode surface area    -   t=elapsed time since application of excitation    -   D=diffusion coefficient    -   C=glucose concentration    -   α=a cofactor-dependent constant.        All of the parameters of this equation will normally be constant        for the sensor except the glucose concentration and time. We can        therefore define a normalized Cottrell failsafe ratio(NCFR) as:        $\begin{matrix}        \begin{matrix}        {{NCFR} = \frac{\sum\limits_{k = 1}^{m}I_{k}}{m\quad I_{m}}} \\        {= \frac{\sum\limits_{k = 1}^{m}{\frac{{nFA}\sqrt{D}}{\sqrt{\Pi}}{Ct}_{k}^{\alpha}}}{m\frac{{nFA}\sqrt{D}}{\sqrt{\Pi}}{Ct}_{m}^{\alpha}}} \\        {= \frac{\sum\limits_{k = 1}^{m}t_{k}^{\alpha}}{{mt}_{m}^{\alpha}}} \\        {= {Constant}}        \end{matrix} & \left( {{Equation}\quad 17} \right)        \end{matrix}$

As the time terms in this equation are known and constant for a sensormeasurement, the ratio always yields a constant for Cottrell curves withidentical sample times and intervals. Therefore, the sum of sensorcurrents divided by the last sensor current should yield a constantindependent of glucose concentration. This relationship is used in thepreferred embodiment to detect potentially faulty biosensor responses.

A Current Sum Failsafe can be devised that places a check on theCottrell response of the sensor by summing all of the acquired currentsduring sensor measurement. When the final current is acquired, it ismultiplied by two constants (which may be loaded into the meter at thetime of manufacture or, more preferably, supplied to the meter with eachlot of sensors, such as by a separate code key or by information codedonto the sensor itself). These constants represent the upper and lowerthreshold for allowable NCFR values.

The two products of the constants multiplied by the final current arecompared to the sum of the biosensor currents. The sum of the currentsshould fall between the two products, thereby indicating that the ratioabove was fulfilled, plus or minus a tolerance.

Therefore, the preferred embodiment performs the following check whenthere is a single DC block: $\begin{matrix}{{\left( I_{m} \right)\left( C_{1} \right)} \leq {\sum\limits_{k = 1}^{m}I_{k}} \leq {\left( I_{m} \right)\left( C_{u} \right)}} & \left( {{Equation}\quad 18} \right)\end{matrix}$where C_(u)=upper constant from the Code Key

-   -   C₁=lower constant from the Code Key    -   I_(m)=final biosensor current

Because some embodiments may contain two DC blocks in the measurementsequence, a Modified Cottrell Failsafe Ratio (MCFR) can be formulatedas: $\begin{matrix}{{MCFR} = \frac{{w_{1}{NCFR}_{1}} + {w_{2}{NCFR}_{2}}}{w_{1} + w_{2}}} & \left( {{Equation}\quad 19} \right)\end{matrix}$where w₁, w₂=weighting constants (e.g. from the Code Key) NCFR₁,NCFR₂=the Normalized Cottrell Failsafe Ratios for DC blocks 1 and 2respectively.Therefore, the preferred embodiment performs the following check whenthere are two DC blocks: $\begin{matrix}{{\left( {w_{1} + w_{2}} \right)I_{m_{1}}I_{m_{2}}C_{L}} \leq \left( {{w_{1}I_{m_{2}}{\sum\limits_{k = 1}^{m_{1}}I_{k}}} + {w_{2}I_{m_{1}}{\sum\limits_{k = 1}^{m_{2}}I_{k}}}} \right) \leq {\left( {w_{1} + w_{2}} \right)I_{m_{1}}I_{m_{2}}C_{u}}} & \left( {{Equation}\quad 20} \right)\end{matrix}$where C_(u)=upper constant from the Code Key

-   -   C_(L)=lower constant from the Code Key    -   I_(m1), I_(m2)=final biosensor current in DC blocks 1 and 2

The NCFR (and MCFR) is correlated with hematocrit. As demonstratedhereinabove in Example 3, the AC phase angle is also correlated withhematocrit. It follows then, that the AC phase angle and the NCFR arecorrelated with one another. This relationship holds only if the sensoris unabused. The correlation degrades for an abused sensor.

It is therefore possible to design an equation to analyze the measuredphase angle data to produce a failsafe calculation that will indicate ifan abused sensor is being used. In the preferred embodiment, it waschosen to use the difference between the phase angles measured at twoseparate frequencies in order to make the test more robust to errorscaused by parasitic resistance, etc. Applying the arctangent function todrive the two populations to different asymptotes yields the followingfailsafe equation:FAILSAFE=1000×arc tan[NCFR/(fs ₀ +fs₁(Φ_(10 kHz)−Φ_(20 kHz)))]  (Equation 21)where 1000=scaling factor

-   -   NCFR=Cottrell Failsafe Ratio    -   fs₀=linear regression intercept    -   fs₁=linear regression slope    -   Φ_(10 kHz)=phase angle at 10 kHz    -   Φ_(2 kHz)=phase angle at 20 kHz

Using Equation 21, the intercept term fs₀ can be chosen such that aFAILSAFE value below zero indicates an abused sensor, while a FAILSAFEvalue above zero indicates a non-abused sensor. Those skilled in the artwill recognise that the opposite result could be obtained by choosing adifferent intercept.

Use of Dose Sufficiency Electrodes

As described hereinabove, it has been recognised that accurate samplemeasurement requires adequate coverage of the measurement electrodes bythe sample. Various methods have been used to detect the insufficiencyof the sample volume in the prior art. For example, the Accu-Chek®Advantage® glucose test meter sold by Roche Diagnostics Corporation ofIndianapolis, Ind. warned the user of the possible inadequacy of thesample volume if non-Cotrellian current decay was detected by the singlepair of measurement electrodes. Users were prompted to re-dose the teststrip within a specified time allotment.

The possibility of insufficient sample size has been heightened inrecent years due to the use of capillary fill devices used inconjunction with blood lancing devices designed to minimize pain throughthe requirement of only extremely small sample volumes. If an inadequateamount of sample is drawn into the capillary fill space, then there is apossibility that the measurement electrodes will not be adequatelycovered and the measurement accuracy will be compromised. In order toovercome the problems associated with insufficient samples, variousprior art solutions have been proposed, such as placing an additionalelectrode downstream from the measurement electrodes; or a singlecounter electrode having a sub-element downstream and major elementupstream of a working electrode; or an indicator electrode arranged bothupstream and downstream from a measurement electrode (allowing one tofollow the flow progression of the sample across the working and counterelectrodes or the arrival of the sample at a distance downstream). Theproblem associated with each of these solutions is that they eachincorporate one or the other electrode of the measurement pair incommunication with either the upstream or the downstream indicatorelectrodes to assess the presence of a sufficient volume of sample toavoid biased test results.

Despite these prior art design solutions, failure modes persist whereinthe devices remain prone to misinterpretation of sample sufficiency. Thepresent inventors have determined that such erroneous conclusions arerelated primarily to the distances between a downstream member of ameasurement electrode pair (co-planar or opposing geometries) and thedose detection electrode, in combination with the diversity ofnon-uniform flow fronts. A sample traversing the capillary fill spacehaving an aberrant (uneven) flow front can close the circuit between ameasurement electrode and an indicator electrode and erroneously advisethe system that sufficient sample is present to avoid a biasedmeasurement result.

Many factors employed in the composition and/or fabrication of the teststrip capillary fill spaces influence such irregular flow frontbehavior. These factors include:

-   -   disparities between surface energies of different walls forming        the capillary fill space.    -   contamination of materials or finished goods in the test strip        manufacturing facility.    -   unintentional introduction of a contaminant from a single        component making up the walls of the capillary fill space (an        example being a release agent (typically silicon) that is common        to manufacturing processes wherein release liners are used).    -   hydrophobic properties of adhesives (or contaminated adhesives)        used in the lamination processes.    -   disparate surface roughnesses on the walls of the capillary fill        space.    -   dimensional aspect ratios.    -   contaminated mesh materials within the capillary fill space.    -   non-homogeneous application of surfactants onto mesh materials        within the capillary fill space.

Another problem with prior art dose sufficiency methodologies determinedby the present inventors relates to the use of one or the other of theavailable measurement electrodes in electrical communication with anupstream or downstream dose detection electrode. In such arrangements,the stoichiometry of the measurement zone (the area above or between themeasurement electrodes) is perturbed during the dose detect/dosesufficiency test cycle prior to making a measurement of the analyte ofinterest residing in the measurement zone. As sample matrices varyradically in make-up, the fill properties of these samples also vary,resulting in timing differences between sample types. Such erratictiming routines act as an additional source of imprecision and expandedtotal system error metrics.

Trying to solve one or more of these obstacles typically can lead to 1)more complex manufacturing processes (additional process steps eachbringing an additional propensity for contamination); 2) additional rawmaterial quality control procedures; 3) more costly raw materials suchas laminate composites having mixtures of hydrophobic and hydrophyllicresins and negatively impacting manufacturing costs; and 4)labor-intensive surfactant coatings of meshes and or capillary walls.

EXAMPLE 8 Determination of Fluid Flow Front Behavior in a Capillary FillSpace

In order to design an electrode system that will adequately indicatedose sufficiency in a test strip employing a capillary fill space, anexperiment was performed to examine the flow front shape at the leadingedge of the sample as it progresses through the capillary fill space.Test fixtures comprising two sheets of clear polycarbonate sheets joinedtogether with double-sided adhesive tape were used, where the capillaryfill space was formed by cutting a channel in the double-sided tape. Useof the polycarbonate upper and lower sheets allowed the flow fronts ofthe sample to be videotaped as it flowed through the capillary fillspace.

Specifically, the test devices were laminated using laser cut 1 mm thickLexan® polycarbonate sheets (obtained from Cadillac Plastics Ltd.,Westlea, Swindon SN5 7EX, United Kingdom). The top and bottompolycarbonate sheets were coupled together using double-sided adhesivetapes (#200MP High Performance acrylic adhesive obtained from 3MCorporation, St. Paul, Minn.). The capillary channels were defined bylaser cutting the required width openings into the double-sided tape.Tape thicknesses of 0.05 μm, 0.125 μm, and 0.225 μm were used to givethe required channel heights. The dimensions of the capillary spaces ofthe test devices are tabulated in FIG. 31.

The top and bottom polycarbonate parts were laminated together with thelaser cut adhesive tapes using a custom-built jig to ensure reproduciblefabrication. For each test device, a fluid receptor region defineing theentrance to the capillary channel was formed by an opening pre-cut intothe upper polycarbonate sheet and adhesive tape components. For each ofthe three channel heights, channel widths of 0.5 mm, 1.00 mm, 1.5 mm,2.00 mm, 3.00 mm, and 4.00 mm were fabricated. The capillary channellength for all devices was 50 mm. Twenty-eight (28) of each of theeighteen (18) device types were constructed. The assembled devices wereplasma treated by Weidman Plastics Technology of Dortmund, Germany. Thefollowing plasma treatment conditions were used:

-   -   Processor: Microwave plasma processor 400    -   Microwave Power: 600W    -   Gas: O₂    -   Pressure: 0.39 miilibar    -   Gas Flow: 150 ml/min    -   Time: 10 minutes    -   Surface Energy Pre-Treatment: <38 mN/m    -   Surface Energy Post-Treatment: 72 mN/m

The plasma-treated devices were stored at 2-8° C. when not in use. Thedevices were allowed to equilibrate to room temperature for one (1) hourminimum before use.

Each of the test devices was dosed with a fixed volume of venous bloodhaving a hematocrit value of 45%. Flow and flow front behavior wascaptured on videotape for later analysis. It was determined that therelative dimensions of the capillary fill channel determined the flowfront behavior. Devices to the left of the dashed line in FIG. 31(devices A2, A4, B2, B4, B5, C2, C4, and C5) resulted in a convex flowfront behavior, while devices to the right of the dashed line (devicesA6, A8, A11, B6, B8, B11, C6, C8, and C11) displayed a concave flowfront behavior. Both the convex and concave flow front behaviors areschematically illustrated in FIG. 31. This data shows that the aspectratio between the height and the width of the capillary fill space is adetermining factor in whether the sample flow front is convex orconcave.

Use of Dose Sufficiency Electrodes cont'd

The problems associated with a concave flow front in a capillary fillspace are illustrated in FIGS. 32A-C. In each of the figures, the teststrip includes a working electrode 3200, a reference electrode 3202, anda downstream dose sufficiency electrode 3204 that works in conjunctionwith one of the measurement electrodes 3200 or 3202. In addition to themeasurement zone stoichiometry problems associated with the use of thedose sufficiency electrode 3204 in conjunction with one of themeasurement electrodes discussed above, FIGS. 32A-C illustrate that asample flow front exhibiting a concave shape can also cause biasedmeasurement results. In each drawing, the direction of sample travel isshown by the arrow. In FIG. 32A, the portions of the sample adjacent tothe capillary walls have reached the dose sufficiency electrode 3204,thereby electrically completing the DC circuit between this electrodeand one of the measurement electrode pair that is being monitored by thetest meter in order to make the dose sufficiency determination. Althoughthe test meter will conclude that there is sufficient sample to make ameasurement at this time, the sample clearly has barely reached thereference electrode 3202 and any measurement results obtained at thistime will be highly biased.

Similarly, FIG. 32B illustrates the situation where the dose sufficiencyelectrode 3204 has been contacted (indicating that the measurementshould be started), but the reference electrode 3202 is only partiallycovered by the sample. Although the sample has reached the referenceelectrode 3202 at this time, the reference electrode 3202 is notcompletely covered by sample, therefore any measurement results obtainedat this time will be partially biased. Both of the situationsillustrated in FIGS. 32A-B will therefore indicate a false positive fordose sufficiency, thereby biasing the measurement test results. Only inthe situation illustrated in FIG. 32C, where the reference electrode3202 is completely covered by the sample, will the measurement resultsbe unbiased due to the extent of capillary fill in the measurement zone.

The present invention solves the stoichiometric problems associated withthe prior art designs pairing the dose sufficiency electrode with one ofthe measurement electrodes when making the dose sufficiencydetermination. As shown in FIG. 33, the present invention comprehends atest strip having an independent pair of dose sufficiency electrodespositioned downstream from the measurement electrodes. The test strip isindicated generally as 3300, and includes a measurement electrode pairconsisting of a counter electrode 3302 and a working electrode 3304. Theelectrodes may be formed upon any suitable substrate in a multilayertest strip configuration as is known in the art and describedhereinabove. The multilayer configuration of the test strip provides forthe formation of a capillary fill space 3306, also as known in the art.Within the capillary fill space 3306, and downstream (relative to thedirection of sample flow) from the measurement electrodes 3302 and 3304are formed a dose sufficiency working electrode 3308 and a dosesufficiency counter electrode 3310, together forming a dose sufficiencyelectrode pair.

When the test strip 3300 is inserted into the test meter, the test meterwill continuously check for a conduction path between the dosesufficiency electrodes 3308 and 3310 in order to determine when thesample has migrated to this region of the capillary fill space. Once thesample has reached this level, the test meter may be programmed toconclude that the measurement electrodes are covered with sample and thesample measurement sequence may be begun. It will be appreciated that,unlike as required with prior art designs, no voltage or current need beapplied to either of the measurement electrodes 3302 and 3304 during thedose sufficiency test using the test strip design of FIG. 33. Thus thestoichiometry of the measurement zone is not perturbed during the dosesufficiency test cycle prior to making a measurement of the analyte ofinterest residing in the measurement zone. This represents a significantimprovement over the dose sufficiency test methodologies of the priorart.

The test strip 3300 is also desirable for judging dose sufficiency whenthe capillary fill space is designed to produce samples that exhibit aconvex flow front while filling the capillary fill space 3306, asillustrated in FIG. 34A. As can be seen, the measurement zone above themeasurement electrodes 3302 and 3304 is covered with sample when theconvex flow front reaches the dose sufficiency electrode pair 3308,3310.The test strip design 3300 may not, however, produce ideal results ifthe capillary fill space 3306 allows the sample to exhibit a concaveflow front while filling, as shown in FIG. 34B. As can be seen, theperipheral edges of the concave flow front reach the dose sufficiencyelectrodes 3308,3310 before the measurement zone has been completelycovered with sample. With DC or low frequency excitation (discussed ingreater detail hereinbelow), the dose sufficiency electrodes 3308,3310will indicate sample sufficiency as soon as they are both touched by theedges of the flow front. Therefore, the dose sufficiency electrodedesign shown in the test strip of FIG. 33 works best when the samplefilling the capillary space 3306 exhibits a convex flow front.

It will be appreciated that the dose sufficiency electrodes 3308,3310have their longest axis within the capillary fill space 3306 orientedperpendicular to the longitudinal axis of the capillary fill space 3306.Such electrodes are referred to herein as “perpendicular dosesufficiency electrodes.” An alternative dose sufficiency electrodearrangement is illustrated in FIGS. 35A-B. As shown in FIG. 35A, thepresent invention also comprehends a test strip having an independentpair of dose sufficiency electrodes positioned downstream from themeasurement electrodes, where the dose sufficiency electrodes have theirlongest axis within the capillary fill space oriented parallel to thelongitudinal axis of the capillary fill space. Such electrodes arereferred to herein as “parallel dose sufficiency electrodes.” The teststrip in FIG. 35 is indicated generally as 3500, and includes ameasurement electrode pair consisting of a counter electrode 3502 and aworking electrode 3504. The electrodes may be formed upon any suitablesubstrate in a multilayer test strip configuration as is known in theart and described hereinabove. The multilayer configuration of the teststrip provides for the formation of a capillary fill space 3506, also asknown in the art. Within the capillary fill space 3506, and downstream(relative to the direction of sample flow) from the measurementelectrodes 3502 and 3504 are formed a dose sufficiency working electrode3508 and a dose sufficiency counter electrode 3510, together forming aparallel dose sufficiency electrode pair.

When the test strip 3500 is inserted into the test meter, the test meterwill continuously check for a conduction path between the dosesufficiency electrodes 3508 and 3510 in order to determine when thesample has migrated to this region of the capillary fill space. Once thesample has reached this level, the test meter may be programmed toconclude that the measurement electrodes are covered with sample and thesample measurement sequence may be begun. It will be appreciated that,as with the test strip 3300 (and unlike as required with prior artdesigns), no voltage or current need be applied to either of themeasurement electrodes 3502 and 3504 during the dose sufficiency testusing the test strip design of FIG. 35. Thus the stoichiometry of themeasurement zone is not perturbed during the dose sufficiency test cycleprior to making a measurement of the analyte of interest residing in themeasurement zone. This represents a significant improvement over thedose sufficiency test methodologies of the prior art.

A further improved operation is realized with the parallel dosesufficiency electrodes of the test strip 3500 when the dose sufficiencyelectrodes are energized with a relatively high frequency AC excitationsignal. When a relatively high frequency AC signal is used as the dosesufficiency excitation signal, the dose sufficiency electrodes 3508,3510display significant edge effects, wherein the excitation signaltraverses the gap between the electrodes only when the electrode edgesalong the gap are covered with the sample fluid. The test strip 3500 isillustrated in enlarged size in FIG. 36 (with only the electrodeportions lying within the capillary fill space 3506 and thestrip-to-meter electrode contact pads visible). When one of the pair ofdose sufficiency electrodes 3508,3510 is excited with an AC signal, themajority of the signal travels from one electrode edge to the edge ofthe other electrode (when the edges are covered with sample), ratherthan from the upper flat surface of one electrode to the upper flatsurface of the other electrode. These paths of edge-to-edge electricalcommunication are illustrated schematically as the electric field lines3602 in FIG. 36.

Higher AC frequencies produce the best edge-only sensitivity from thedose sufficiency electrodes. In the preferred embodiment, a 9 mV_(rms)(±12.7 mV peak-to-peak) excitation signal of 10 kHz is used to exciteone of the dose sufficiency electrodes. The gap width GW between theedges of the dose sufficiency electrodes 3508,3510 is preferably 100-300μm, more preferably 150-260 μm, and most preferably 255 μm. A smallergap width GW increases the amount of signal transmitted between dosesufficiency electrodes whose edges are at least partially covered bysample; however, the capacitance of the signal transmission pathincreases with decreasing gap width GW.

An advantage of the parallel dose sufficiency electrode design of FIGS.35 and 36, when used with AC excitation, is that there is substantiallyno electrical communication between the electrodes until the samplecovers at least a portion of the edges along the electrode gap.Therefore, a sample exhibiting the concave flow front of FIG. 35A, wherethe illustrated sample is touching both of the dose sufficiencyelectrodes 3508,3510 but is not touching the electrode edges along thegap, will not produce any significant electrical communication betweenthe dose sufficiency electrodes. The test meter will therefore not forma conclusion of dose sufficiency until the sample has actually bridgedthe dose sufficiency electrodes between the electrode edges along thegap. This will happen only after the rear-most portion of the concaveflow front has reached the dose sufficiency electrodes 3508,3510, atwhich point the sample has completely covered the measurement zone overthe measurement electrodes. As can be seen in FIG. 35B, convex sampleflow fronts will activate the dose sufficiency electrodes 3508,3510 assoon as the flow front reaches the dose sufficiency electrodes (at whichpoint the sample has completely covered the measurement zone over themeasurement electrodes).

Another advantage to the parallel dose sufficiency electrodesillustrated in FIGS. 35 and 36 is that the amount of signal transmittedbetween the electrodes is proportional to the amount of the gap edgesthat is covered by the sample. By employing an appropriate thresholdvalue in the test meter, a conclusion of dose sufficiency can thereforebe withheld until the sample has covered a predetermined portion of thedose sufficiency electrode gap edge. Furthermore, an analysis of thedose sufficiency signal will allow the test meter to record thepercentage of fill of the capillary fill space for each measurement madeby the test meter, if desired.

While the electrode geometry itself demonstrates an advantage overprevious embodiments in terms of detecting an adequate sample,particularly in the case of a convex flow front, it was found thatfurther improvement is achieved in the use of AC responses over DCresponses for sample detection. DC responses have the problems of beingsensitive to variations in, for example, temperature, hematocrit and theanalyte (glucose for example). AC responses at sufficiently highfrequency can be made robust to the variation in the analyteconcentration. Further, the AC response generated at sufficiently highfrequencies in such capillary fill devices is primarily limited by theamount of the parallel gap between the electrode edges which is filledby the sample. Thus, for a convex flow front, little or no AC response(in this case admittance) is perceived until the trough of the flowfront actually intrudes within the parallel edges of the samplesufficiency electrodes. Further, by means of threshold calibration, thesensor can be made more or less sensitive as is deemed advantageous,with a higher threshold for admittance requiring more of the parallelgap to be filled before test initiation.

A further limitation of existing devices is the inability of theelectrode geometry to discern the amount of time needed to fill thecapillary space of the sensor. This limitation is caused by havinginterdependence of the dose sufficiency electrode and the measurementelectrodes. This is a further advantage of independent dose sufficiencyelectrodes. In the preferred embodiment a signal is first applied acrossthe measurement electrodes prior to dosing. When a response is observed,the potential is immediately switched off and a second signal is appliedacross the dose sufficiency electrodes during which time the system bothlooks for a response to the signal (indicating electrode coverage) andmarks the duration between the first event (when a response is observedat the measurement electrodes) and the second event (when a response isobserved at the dose sufficiency electrodes). In cases where very longintervals may lead to erroneous results, it is possible to establish athreshold within which acceptable results may be obtained and outside ofwhich a failsafe is triggered, preventing a response or at a minimumwarning the user of potential inaccuracy. The amount of time lag betweendosing and detection of a sufficient sample that is considered allowableis dependent upon the particular sensor design and chemistry.Alternatively, an independent pair of dose detection electrodes (notshown) may be added upstream from the measurement electrodes in order todetect when the sample is first applied to the sensor.

While a DC signal could be used for detection in either or both of theabove events, the preferred embodiment uses an AC signal at sufficientlyhigh frequency to avoid unnecessarily perturbing the electrochemicalresponse at the measurement electrodes and to provide robust detectionwith respect to flow front irregularities.

All publications, prior applications, and other documents cited hereinare hereby incorporated by reference in their entirety as if each hadbeen individually incorporated by reference and fully set forth.

While the invention has been illustrated and described in detail in thedrawings and foregoing description, the description is to be consideredas illustrative and not restrictive in character. Only the preferredembodiment, and certain other embodiments deemed helpful in furtherexplaining how to make or use the preferred embodiment, have been shown.All changes and modifications that come within the spirit of theinvention are desired to be protected.

1. A meter for use in combination with a disposable electrochemicalsensor for detection and/or quantification of an analyte in a liquidsample comprising a timing circuit for controlling the measurement ofcurrent indicative of analyte in the sample following detection ofsample application to a test strip inserted in the meter, wherein thetiming circuit causes the measurement of current to occur at a time 10seconds or less after the detection of sample application.